Coordinating operation of multiple lower limb devices

ABSTRACT

Systems and methods for operating autonomous lower limb devices, including at least one adjustable joint and configured for at least supporting a body are provided. A method includes collecting real-time sensor information associated with the autonomous lower limb device. The method also includes receiving, over a communications link, remote data for at least one other lower limb device configured for supporting the body. The method further includes generating control data for transitioning the autonomous lower limb device from a current state in a current finite state model for operating the autonomous lower limb device to a different state in the current finite state model, where the different state is selected based on the real-time sensor information and the remote data.

This application is a continuation of U.S. Non-Provisional application Ser. No. 13/826,069, filed Mar. 14, 2013, now U.S. Pat. No. 9,289,317, and entitled “COORDINATING OPERATION OF MULTIPLE LOWER LIMB DEVICES”, which claims priority to and the benefit of U.S. Provisional Patent Application No. 61/610,869, filed Mar. 14, 2012 and entitled “COMMUNICATION NETWORK FOR MULTIPLE PROSTHETIC DEVICES”.

FIELD OF THE INVENTION

The present invention relates to the operation of multiple lower limb devices used for at least supporting a body, and more specifically to apparatus and methods for coordinating to the operation of multiple lower limb devices used for at least supporting a body.

BACKGROUND

Bilateral amputees constitute a relatively small proportion of lower limb amputees. Despite this, a bilateral amputation is a much more significant disability than a unilateral amputation, particularly in the case of transfemoral amputation. Specifically, in the case of unilateral amputation, an amputee is able to compensate substantially with his or her sound side to address biomechanical deficiencies in the prosthetic limb. For example, relative to healthy individuals, unilateral transfemoral amputees rely disproportionately on their sound side leg for providing the net power output required for stair ascent, slope ascent, and for sit-to-stand transitions, all of which require net positive power at the knee and/or ankle joints. Unilateral transfemoral amputees also provide significant compensatory effort with their sound side leg in activities which do not necessarily require net positive power, such as level walking, slope and stair descent, standing (particularly on uneven terrain), and stand-to-sit transitions. For example, “heel hiking” is a common compensatory action observed during sound-side stance in level walking. Heel hiking elevates the amputee's center of mass, in order to increase the swing-side clearance between the prosthetic foot and ground, in order to decrease the likelihood of scuffing or stumbling with the prosthetic leg.

Unlike unilateral amputees, bilateral transfemoral amputees are not able to compensate with a sound-side limb for deficiencies in gait biomechanics on the prosthetic side, and in particular are not able to compensate similarly for deficiencies in power generation.

SUMMARY

Embodiments of the invention concern systems and methods for coordinating the operation of multiple, autonomous lower limb devices used for at least supporting a body. In a first embodiment, there is provided a method of operating a plurality of autonomous lower limb devices configured for at least supporting a body and that each include at least one adjustable joint. The method includes, at a first lower limb device of the plurality of the autonomous lower limb devices, collecting real-time sensor information associated with the first lower limb device. The method also includes receiving at the first lower limb device remote data for at least a second lower limb device of the plurality of autonomous lower limb devices. The method further includes generating control data for transitioning the at least one adjustable joint of the first lower limb device from a current state in a finite state model for operating the first lower limb device to a different state in the finite state model, where the different state is selected based on the real-time sensor information and the remote data. In the method, the remote data can consist essentially of at least one of real-time sensor information for the second lower limb device or state data associated with the second lower limb device.

In the method, the generating can include configuring the control data to transition the first lower limb device to a non-weight bearing state only if the state data in the remote data indicates a weight bearing state. The generating can also include determining a local terrain configuration based on the real-time input and the remote data and configuring the control data based on the local terrain. The generating can further include detecting a perturbation based on the remote data and configuring the control data to compensate for the perturbation.

In the method, responsive to a failure in the communications link, the generating can include configuring the control data in accordance with pre-defined set of safe states. The generating can also include detecting one of a sit-to-stand transition and a stand-to-sit transition based on the real-time input and the remote data to yield a detected transition and configuring the control data based on the detected transition.

In a second embodiment of the invention, there is provided a computer-readable medium having stored thereon a plurality of instructions for causing a controller device in each of a plurality of autonomous lower limb devices associated with a body to perform any of the methods described above.

In a third embodiment of the invention, there is provided an autonomous lower limb device for at least supporting a body. The lower limb device can include at least one adjustable joint and at least one sensor for collecting real-time sensor information for the lower limb device. The lower limb device can also include a communications device and a processor communicatively coupled to the at least one sensor, the adjustable joint, and the communications device. The lower limb device can further include a computer-readable medium, having stored thereon instructions for causing the processor to perform the steps of generating first control data for the communications device to establish a communications link with at least one other lower limb device configured for at least supporting the body, receiving, over the communications link, remote data for the at least one other lower limb device, and generating second control data for transitioning the lower limb device from a current state in finite state model for operating the lower limb device to a different state in the finite state model, where the different state is selected based on the real-time sensor information and the remote data. In the lower limb device, the remote data can consist essentially of at least one of real-time sensor information for the other device or state data associated with the other device.

Further, the instructions for generating the first control data can further include instructions for causing the processor to configure the first control data to transition the lower limb device to a non-weight bearing state only if the state data in the remote data indicates a weight bearing state. The instructions for generating the first control data can also include instructions for causing the processor to determine a local terrain configuration based on the real-time input and the remote data and configure the first control data based on the local terrain. The instructions for generating the first control data also can further include instructions for causing the processor to detect a perturbation based on the remote data, and configure the first control data to compensate for the perturbation. The instructions also can include instructions for causing the processor to detect a failure in the communications link and the instructions for generating the first control data can further include instructions for causing the processor to configure the first control data in accordance with pre-defined set of safe states in response to the failure. The instructions for generating the first control data can also include instructions for causing the processor to detect one of a sit-to-stand transition and a stand-to-sit transition based on the real-time input and the remote data to yield a detected transition, and configure the first control data based on the detected transition.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a view of a powered knee and ankle prosthesis, according to an embodiment of the invention;

FIG. 1B is an exploded view of the powered knee and ankle prosthesis shown in FIG. 1A, according to an embodiment of the invention;

FIG. 2 is an exploded view of knee motor unit, according to an embodiment of the invention;

FIG. 3 is an exploded view of ankle motor unit, according to an embodiment of the invention;

FIG. 4 is an exploded view of knee joint, according to an embodiment of the invention;

FIG. 5 is an exploded view of ankle joint, according to an embodiment of the invention;

FIGS. 6A and B are views of a foot having toe and heel force sensing elements, according to an embodiment of the invention;

FIG. 7 shows the joint angle and torque convention used herein. Positive torque is defined in the direction of increasing angle;

FIG. 8 shows the subdivision of normal walking into four internal phases showing the knee and ankle angles during the phases, according to an embodiment of the invention;

FIG. 9 shows a finite-state model of normal walking, according to an embodiment of the invention. Each box represents a different internal phase and the transition conditions between the internal phases are specified;

FIG. 10 shows piecewise fitting of knee and ankle torques during normal speed level walk scaled for a 75 kg adult to a non-linear spring-damper impedance model;

FIG. 11 is a diagram for an active/passive decomposition based control of the powered knee and ankle prosthesis, according to an embodiment of the invention;

FIG. 12 is a diagram for a general form of active-passive decomposition control including intent recognition that provides supervisory modulation, according to an embodiment of the invention;

FIG. 13A is a side view of powered knee and ankle prosthesis, according to another embodiment of the invention;

FIG. 13B is a front view of powered knee and ankle prosthesis of FIG. 13A.

FIGS. 14A and 14B show perspective and bottom views of an exemplary sagittal moment load cell suitable for use in the various embodiments of the invention;

FIG. 15 is a block diagram of an exemplary embedded microcontroller in accordance with an embodiment of the invention;

FIG. 16 is a control state chart for the three activity modes corresponding to walking, standing, and sitting, and for the internal phases and their corresponding transitions within each activity mode;

FIG. 17 shows knee angle modulated knee stiffness during pre-stand (solid line) and pre-sit (dashed line) phases;

FIG. 18 is a plot of axial actuation unit force versus ankle angle;

FIG. 19 shows a normal speed walking phase portrait of the knee joint and four stride segments;

FIG. 20 shows the selection of indexing data samples during a first segment of a walking stride;

FIGS. 21A, 22B, and 22C are the output of the decomposition for Segment 1 showing the spring and dashpot constants and the active and passive knee torques;

FIG. 22 is a state chart for governing the discrete dynamics of an active-passive decomposition controller in accordance with an embodiment of the invention;

FIG. 23 is a state chart for governing the discrete dynamics of the cadence estimator in accordance with an embodiment of the invention;

FIG. 24 is a schematic diagram of accelerometer measurements for slope estimation in accordance with an embodiment of the invention;

FIG. 25 is a state chart for slope estimation in a controller in accordance with an embodiment of the invention;

FIGS. 26A and 26B show front and back views of a friction/cable drive motor in accordance with an embodiment of the invention;

FIG. 27 shows an exemplary embodiment of a belt drive transmission in accordance with an embodiment of the invention;

FIGS. 28A and 28B show side views of first and second positions, respectively, achievable for an exemplary embodiment of a chain drive transmission including an eccentric mount in accordance with an embodiment of the invention;

FIG. 29 illustrates schematically the components for the adjustable bearing mounts in FIGS. 28A and 28B;

FIG. 30 illustrates an exemplary configuration of a powered leg prosthesis in accordance with the embodiments shown in FIGS. 27-29;

FIG. 31 shows exemplary detailed state chart for an autonomous powered leg prosthesis in accordance with the various embodiments;

FIG. 32 shows exemplary state charts for bilateral, autonomous powered leg prostheses linked via a controller area network in accordance with the various embodiments; and

FIG. 33 shows plots of hip angle, knee angle, ankle angle, as a percentage of stride, for healthy subjects, subjects utilizing bilateral powered prostheses configured in accordance with the various embodiments, and subjects utilizing bilateral passive prostheses.

DETAILED DESCRIPTION

The present invention is described with reference to the attached figures, wherein like reference numerals are used throughout the figures to designate similar or equivalent elements. The figures are not drawn to scale and they are provided merely to illustrate the instant invention. Several aspects of the invention are described below with reference to example applications for illustration. It should be understood that numerous specific details, relationships, and methods are set forth to provide a full understanding of the invention. One having ordinary skill in the relevant art, however, will readily recognize that the invention can be practiced without one or more of the specific details or with other methods. In other instances, well-known structures or operations are not shown in detail to avoid obscuring the invention. The present invention is not limited by the illustrated ordering of acts or events, as some acts may occur in different orders and/or concurrently with other acts or events. Furthermore, not all illustrated acts or events are required to implement a methodology in accordance with the present invention.

Recently, a number of powered lower limb prostheses have begun to emerge. The authors have previously developed a powered transfemoral prosthesis, and demonstrated in unilateral amputees improved kinetics for a variety of activities, as described in U.S. patent application Ser. No. 13/537,530 filed Jun. 29, 2012, the contents of which are herein incorporated by reference in their entirety. Like the biomechanically intact lower limb, such prostheses are capable of generating net positive power at the knee and/or ankle joints. Recent publications indicate that such prostheses can provide improved kinematics, energetics and stability for unilateral lower limb amputees.

In view of the limitations of existing lower limb prostheses, the various embodiments of the invention provide a needed system and method using multiple controlled lower limb devices for at least supporting a body. As used herein, the term “body” refers to a physical system, whether biological (e.g., amputee) or non-biological, that operates independently of the lower limb devices, but through which the lower limb devices are physically connected so as to support and, optionally, propel, the physical system. For example, the present inventors have recognized that if a first controlled leg prosthesis is able to receive or discern information regarding the configuration of a contralateral leg of a patient including a second controlled leg prosthesis, the first controlled leg prosthesis can be controlled to improve the stability and safety of the patient. In the various embodiments, this can be provided by allowing the separately controlled lower limb devices to communicate configuration, state, or mode information. Thereafter, based on such information, each of these lower limb device can independently determine a proper state to operate in. In other words, the lower limb devices collectively achieve greater awareness than is possible with isolated, independently operating lower limb devices.

Although the description below will be primarily directed to powered leg prostheses for bipedal motion for bilateral amputees with transfemoral prostheses, this is for exemplary purposes only. The various embodiments are equally applicable to amputee with one prosthesis that is a powered knee and ankle prosthesis for a transfemoral amputation and another prosthesis is a powered ankle prosthesis for a transtibial amputation. Further, the various embodiments are also useful for powered leg prostheses for bipedal motion for bilateral amputees with transtibial prostheses. Additionally, the various embodiments are equally applicable to any type of lower limb devices, including any other type of prostheorthotic and robotic systems with any number of jointed devices. Further, the various embodiments are equally applicable to any type of powered or passive lower limb devices

Prior to discussing the coordinating of multiple lower limb devices, the disclosure first turns to FIGS. 1A-30 where there are described various configurations for powered leg and ankle prostheses, including a controller, which can be modified to include a running controller in accordance with the various embodiments of the invention.

Exemplary Prosthesis Configurations

A first design for a prosthesis for use in the various embodiments of the invention is shown in FIG. 1A through FIG. 6B. The prosthesis 100 comprises a prosthetic lower leg 101. Lower leg 101 can be coupled to a powered knee joint comprising a knee motor unit 105 coupled to a knee joint 110, and a powered ankle joint comprising an ankle motor 115 coupled to an ankle joint 120. A sagittal plane moment sensor 125 can be located between the prosthesis and the user to measure the moment, and in one embodiment is located immediately below the socket interface. In the embodiment shown, sensor 125 measures the sagittal plane moment, while separate sensors described below measure the ball of foot force and heel force with respect to the ground or other object the foot is pressed against. A load sensor 135 can be positioned at the ball of the foot, and a load sensor 140 can be positioned at the heel of the foot. However, in another embodiment (not shown) sensor 125 can measure the sagittal plane moment, the frontal plane moment and the axial force, such as provided by the three-axis socket load cell. This alternate embodiment can eliminate the need for sensor 135 and sensor 140.

Load sensors 141 and 142 are in series with each motor unit 105 and 115, respectively for motor unit force control. Position sensors 151 and 152 are provided at each joint 110 and 120 as shown in FIGS. 4 and 5 respectively. Position sensors 151 measure joint angle (0 as used below) and can be embodied as potentiometers. The computer/process controller, power source (e.g. a battery such as a Li ion battery, and electrical connections in the case of an electrical power source are not shown to avoid obscuring aspects of the invention. Non-electrical power sources may also be used, such as pneumatic power, or non-battery electrical sources, such as hydrogen-based fuel cells.

Prosthesis 100 is shown in an exploded view in FIG. 1B. Joints 110 and 120 are more clearly shown as compared to FIG. 1A.

FIG. 2 is an exploded view of knee motor unit 105, according to an embodiment of the invention. Load sensor 141 is shown as a load cell (e.g. strain gauge). Load sensor 141 measures force and moments. The motor unit 105 comprises a motor-driven ball screw assembly which drives the knee joint through a slider-crank linkage comprising screw 145. Other motor drive assemblies may also generally be used.

FIG. 3 is an exploded view of ankle motor unit 115, according to an embodiment of the invention. Load sensor 142 is generally analogous to load sensor 141. The motor unit 115 comprises a motor-driven ball screw assembly which drives the ankle joint through a slider-crank linkage comprising screw 145. The ankle motor 115 includes a spring 147 positioned to provide power in parallel (thus being additive) with power provided by the motor unit 115. Spring 147 biases the motor unit's force output toward ankle plantarflexion, and supplements the power output provided by motor unit 115 during ankle push off.

FIG. 4 is an exploded view of knee joint 110, according to an embodiment of the invention. As described above, knee joint 110 includes position sensor 151 that can be embodied as a potentiometer for angle measurements of the knee joint 110.

FIG. 5 is an exploded view of ankle joint 120, according to an embodiment of the invention. As described above, ankle joint 120 includes position sensor 152 that can be embodied as a potentiometer for angle measurements of the ankle joint 120.

FIG. 6A is a view of a foot 170 having ball of foot sensors 135, according to an embodiment of the invention. Sensors 135 are provided to measure the ground reaction forces near the ball of the foot, such as when the foot strikes the ground. FIG. 6B is a view of a foot 170 having ball of foot sensors 135 and heel sensors 140, according to an embodiment of the invention. Sensors 140 are provided to measure the ground reaction forces on the heel of the foot when the foot 170 strikes the ground. Sensors 135 and 140 can be embodied as strain based sensors.

Unlike existing passive prostheses, the introduction of power into a prosthesis according to embodiments of the invention provides the ability for the device to also act, rather than simply react. As such, the development of a suitable controller and control methodology that provides for stable and reliable interaction between the user and prosthesis is provided herein. Control according to embodiments of the invention has been found to enable the user to interact with the prosthesis by leveraging its dynamics in a manner similar to normal gait, and also generates more stable and more predictable behavior.

Thus, rather than gather user intent from the joint angle measurements from the contralateral unaffected leg, embodiments of the invention infer commands from the user via the (ipsilateral) forces and moments of interaction between the user and prosthesis. Specifically, the user interacts with the prosthesis by imparting forces and moments from the residual limb to the prosthesis, all of which can be measured via suitable sensor(s), such as sensors 125, 140 and 141 described above which measures moments/forces. These forces and moments serve not only as a means of physical interaction, but also serve as an implicit communication channel between the user and device, with the user's intent encoded in the measurements. Inferring the user's intent from the measured forces and moments of interaction according to embodiments of the invention provides several advantages relative to the known echo approach.

In one embodiment of the invention the torque required at each joint during a single stride (i.e. a single period of gait) can be piecewise represented by a series of passive impedance functions. A regression analysis of gait data indicates that joint torques can be characterized by functions of joint angle (θ) and angular velocity by an impedance model, such as the following exemplary passive impedance function shown in equation 1 below: τ=k ₁(θ−θ_(e))+b*  (1) where k₁, b, and the equilibrium joint angle θ, are all constants that are generally generated empirically, and are constants for a given joint during a given internal phase (e.g. knee, internal phase 3). k₁ characterizes the linear stiffness. b is the linear damping coefficient, θ is the measured joint angle which can characterize the state of the prosthesis, θ_(e) is the equilibrium angle, {dot over (θ)} is the angular velocity of the joint, and τ is the joint torque. Given these constants, together with instantaneous sensor measurements for θ and {dot over (θ)}, the torque (τ) at the joints (knee and ankle) can be determined.

Positive directions of the angle (θ) and torque (τ) as used herein are defined as shown in FIG. 7. If the coefficients b and k₁ are constrained to be positive, then the joints will each exponentially converge to a stable equilibrium at θ=θ_(e) and {dot over (θ)}=θ within each internal phase. That is, within any given internal phase, the actuators are energetically passive (i.e. the joint will come to rest at a local equilibrium). While the unactuated prosthesis can be energetically passive, the behavior of one joint (knee or ankle) or the combined behavior of the knee and ankle joints, can be likewise passive, and thus will generally respond in a predictable manner.

Responsive to direct input from the user (e.g. a heel strike) to trigger a change in internal phase, power (torque) can be delivered from the power source (e.g. battery) to the prosthesis in the proper magnitude to provide the desired movement. Since the switching can be triggered by direct input from the user related to the current internal phase, the user maintains direct influence over the power applied to the prosthesis. If the user does not trigger the next internal phase (i.e. remains stationary) no net energy is delivered. That is, the prosthesis will generally cease to receive power from the power source for moving the joint, and will instead, due to the damped response, soon come to rest at the local equilibrium identified with the present internal phase.

As described above, the decomposition of joint behavior into passive segments requires the division of the gait cycle into a plurality of internal phases or “finite states” characterized by an impedance function and a set of constants for the impedance function, as dictated by their functions and the character of the piecewise segments of the impedance functions described above. The switching rules between internal phases should generally be well defined and measurable, and the number of phases should be sufficient to provide a substantially accurate representation of normal joint function. In one embodiment of the invention, the swing and stance phase of gait can constitute a minimal set of internal phases.

Based on least-squares regression fitting of Equation 1 to empirical gait data, the present Inventors determined that such fits were improved significantly by further dividing the two modes of swing and stance each into two additional internal phases to realize four phases, as shown in FIG. 8. A fifth internal phase can also be added, as illustrated in FIG. 16. The angle (θ) of the prosthetic knee (above) and ankle joint (below) can be provided during each internal phase as a function of the % of the stride. Angle values shown can be used as threshold values to trigger phase changes as described below relative to FIG. 9. As clear to one having ordinary skill in the art, the number of phases can be other than two or four.

FIG. 9 shows exemplary switching rules between internal phases for walking. FIG. 16 shows another set exemplary switching rules, for walking, standing, and sitting activity modes. As described above, if the user does not initiate actions that trigger the next phase (e.g. based on the switching rules), the prosthesis will cease to receive power and will come to rest at the local equilibrium identified with the present phase. For example, switching can be based on the ankle angle>a threshold value (Mode 1 to Mode 2), or ankle torque<threshold) (Mode 2 to Mode 3), the angle or torque measurements provided by on board sensors as described above.

Phase 1 shown in FIG. 8 begins with a heel strike by the user (which can be sensed by the heel force sensor), upon which the knee immediately begins to flex so as to provide impact absorption and begin loading, while the ankle simultaneously plantarflexes to reach a flat foot state. Both knee and ankle joints have relatively high stiffness (and can be accounted for by k1 in equation 1) during this phase to prevent buckling and allow for appropriate stance knee flexion, because phase 1 comprises most of the weight bearing functionality. Phase 2 is the push-off phase and begins as the ankle dorsiflexes beyond a given angle (i.e. user's center of mass lies forward of stance foot). The knee stiffness decreases in this mode to allow knee flexion while the ankle provides a plantarflexive torque for push-off. Phase 3 begins as the foot leaves the ground as detected by the ankle torque load cell and lasts until the knee reaches maximum flexion. Mode 4 is active during the extension of the knee joint (i.e. as the lower leg swings forward), which begins as the knee velocity becomes negative and ends at heel strike (e.g. as determined by the heel force sensor).

In both of the swing phases (Phases 3 and 4), the ankle torque can be small and can be represented in the controller as a (relatively) weak spring regulated to a neutral position. The knee can be primarily treated as a damper in both swing phases.

Impedance modeling of joint torques was preliminarily validated by utilizing the gait data from a healthy 75 kg subject, as derived from body-mass normalized data. Incorporating the four internal phases described above, along with the motion and torque data for each joint, a constrained least-squares optimization was conducted to generate a set of parameters k₁, b and θ_(e) for each phase for each joint for use in Equation 1. The resulting parameter set can be fit to joint torques and is shown graphically in FIG. 10. FIG. 10 shows piecewise fitting of knee and ankle torques during normal speed level walk scaled for a 75 kg adult to a non-linear spring-damper impedance model. The numbers shown in each phase represent the mean ratio of the stiffness forces to damping forces predicted by the fit. The vertical lines represent the segmentation of a gait stride into four distinct phases. The fit shown in FIG. 10 clearly indicates that normal joint function can be represented by the use of piecewise passive functions.

Controllers according to embodiments of the invention generally comprise an underlying gait controller (intra-modal controller). An optional supervisory gait controller (also called intent recognizer) can also be provided. Both controllers generally utilize measured information. This information generally comprises user and ground interaction forces (F) and moments/torques (τ), joint angles and angular velocities from on-board sensors, and can be used to extract real-time input from the user. The gait control component utilizes the sensed instantaneous nature of the user input (i.e., moments and forces) to control the behavior of the leg within a given activity mode, such as standing, walking, or stair climbing.

Two exemplary approaches to intra-modal impedance generation are described below. The first approach is shown in FIG. 11 and represents a general form of active-passive decomposition-based intra-mode control. The second embodiment shown in FIG. 12 includes the control structure shown in FIG. 11 but adds a supervisory intent recognizing controller to modulate the intra-modal control based on inputs from an intent recognition module. As shown in FIGS. 11 and 12, F_(s) is the force the user of the prosthesis is applying, such as a heel force in the case of a heel strike, τ represents joint torque, and θ represent joint angles. τ_(a) represents the active component of joint torque which is roughly proportional to the input force, and τ_(p) represents the passive component of torque. The active joint torque τ_(a) is thus the total joint torque τ minus the passive joint torque, τp. Derivatives are shown using the dot convention, with one dot being the first derivative (e.g., {dot over (θ)} being angular velocity) and two dots representing the second derivative.

In the embodiment of the intra-modal controller shown in FIG. 11, the behavior of the prosthesis can be decomposed into a passive component and an active control component. The active control component is an algebraic function of the user's real-time input F_(s) (i.e., sensed socket-prosthesis interface forces and moments and sensed ground reaction forces). The controller output is shown as the active torque (τ_(a)) minus the passive torque τ_(p). The controller output τ_(a)-τ_(p) applied to the prosthetic leg based on dynamics of the leg responds via θ and {dot over (θ)}. The system response, θ and {dot over (θ)}, is fed back to the controller.

Power applied to the prosthesis can be thus commanded directly by the user through measured interface forces and moments initiated by user movements. In the absence of these commands from the user, F_(s)=0, τ_(a)=0 and the prosthesis fundamentally (by virtue of the control structure) cannot generate power, and thus only exhibits controlled passive behavior. Due to the decomposition of energetic behaviors inherent in this control structure, the prosthesis under it's own control can be generally stable and passive. Unlike known echo control approaches, the input can be real-time, based only on the affected leg, and thus the approach can be equally applicable to bilateral and unilateral amputees and can reflect the instantaneous intent of the user. Additionally, unlike echo control that is based on servocontrol, the prosthesis will exhibit a natural impedance to the user that should feel more like a natural limb. These combined features should result in an active prosthesis that will feel inasmuch as possible like a natural extension of the user. The structure and properties of both the gait controller and intent recognizer are described below.

As described above, since gait is largely a periodic activity, joint behavior can be functionally decomposed over a period by decomposing the joint torque into a passive component and an active component. The passive component can comprise a function of angle (i.e., single-valued and odd), and a function of angular velocity passive (i.e., single-valued and odd), such as equation 1 described above. The active component can be a function of the user input (i.e., socket interface forces). Given a set of data that characterizes a nominal period of joint behavior, the passive component can be first extracted from the whole, since the passive behavior is a subset of the whole (i.e., the passive component consists of single-valued and odd functions, while the active has no restrictions in form). The passive component can be extracted by utilizing a least squares minimization to fit a generalized singled-valued odd function of angle and angular velocity to the torque. Once the passive component is extracted, the residual torque (i.e., the portion that is not extracted as a passive component), can be constructed as an algebraic function of the sensed socket interface and ground reaction forces (i.e., the direct-acting user input) by incorporating a similar candidate function, but not restricted to be of passive form. Finally, superimposing the passive and active components provides a decomposed functional approximation of the original period joint torque.

In the embodiment of the intra-modal controller shown in FIG. 12, a supervisory intent recognizer can be added that utilizes the same sensed user inputs (i.e., moments and forces) as the intra-modal/gait controller, but extracts the user's intent based on the characteristic shape of the user input(s) and system response (e.g. F, θ, θ-dot). Based on the extracted intent, the supervisory intent recognizer modulates the behavior of the underlying gait controller to smoothly transition behavior within a gait (e.g., speed and slope accommodation) and between gaits (e.g., level walk to stair ascent), thus offering a unified control structure within and across all gaits.

Gait intent recognition can be a real-time pattern recognition or signal classification problem. The signal in this case is generally the combination of socket interface forces Fs and the dynamic state of the prosthesis, which in one embodiment can be a vector of the knee and ankle angles θ for a powered leg prosthesis according to an embodiment of the invention. A variety of methods exist for pattern recognition and signal classification including nearest neighbor algorithms, neural networks, fuzzy classifiers, linear discriminant analysis, and genetic algorithms.

As described above, embodiments of the invention include a number of sensors for providing signals for adjusting operation of a leg and ankle prosthesis. A description of one exemplary arrangement of sensors can be described below with respect to FIGS. 13A, 13B, 14A, and 14B. FIG. 13A is a side view of powered knee and ankle prosthesis 1300, according to another embodiment of the invention. FIG. 13B is a front view of powered knee and ankle prosthesis of FIG. 13A. FIGS. 14A and 14B show perspective and bottom views of an exemplary sagittal moment load cell suitable for use in the various embodiments of the invention.

Each joint actuation unit, such as knee actuation unit 1302 and ankle actuation unit 1304 in FIG. 13A, can include a uniaxial load cell positioned in series with the actuation unit for closed loop force control. Both the knee and ankle joints can incorporate integrated potentiometers for joint angle position. The ankle actuation unit can include a spring 1305, as described above with respect to FIGS. 1A-4. One 3-axis accelerometer can be located on the embedded system 1306 and a second one can located below the ankle joint 1308 on the ankle pivot member 1310. A strain based sagittal plane moment sensor 1312, such as the sensor shown in FIGS. 14A and 14B, can located between the knee joint 1314 and the socket connector 1316, which measures the moment between a socket and the prosthesis. In the various embodiments of the invention, a sagittal plane moment sensor can be designed to have a low profile in order to accommodate longer residual limbs. The sensor can incorporate a full bridge of semiconductor strain gages which measure the strains generated by the sagittal plane moment. In one embodiment of the invention, the sagittal plane moment sensor was calibrated for a measurement range of 100 Nm. A custom foot 1318 can designed to measure the ground reaction force components at the ball 1320 of the foot and heel 1322. The foot can include of heel and ball of foot beams, rigidly attached to a central fixture and arranged as cantilever beams with an arch that allows for the load to be localized at the heel and ball of the foot, respectively. Each heel and ball of foot beam can also incorporates a full bridge of semiconductor strain gages that measure the strains resulting from the respective ground contact forces. In one embodiment of the invention, the heel and ball of foot load sensors were calibrated for a measurement range of 1000 N. In addition, incorporating the ground reaction load cell into the structure of a custom foot can eliminate the added weight of a separate load cell, and also enables separate measurement of the heel and ball of foot load. The prosthetic foot can be designed to be housed in a soft prosthetic foot shell (not shown).

The powered prostheses described above contain an embedded microcontroller that allows for either tethered or untethered operation. An exemplary embedded microcontroller system 1500 is shown in the block diagram in FIG. 15. The embedded system 1500 consists of signal processing, power supply, power electronics, communications and computation modules. The system can be powered by a lithium polymer battery with 29.6 V. The signal electronics require +/−12 V and +3.3 V, which are provided via linear regulators to maintain low noise levels. For efficiency, the battery voltage can be reduced by PWM switching amplifiers to +/−15 V and +5 V prior to using the linear regulators. The power can be disconnected via a microcontroller that controls a solid state relay. The power status can be indicated by LED status indicators controlled also by the microcontroller.

The analog sensor signals acquired by the embedded system include the prosthesis sensors signals (five strain gage signals and two potentiometer signals), analog reference signals from the laptop computer used for tethered operation, and signals measured on the board including battery current and voltage, knee and ankle servo amplifier currents and two 3-axis accelerometers. The prosthesis sensor signals are conditioned using input instrumentation amplifiers. The battery, knee motor and ankle motor currents are measured by current sense resistors and current sensing amplifiers. The signals are filtered with a first-order RC filter and buffered with high slew rate operational amplifiers before the analog to digital conversion stage. Analog to digital conversion can be accomplished by two 8-channel analog to digital convertors. The analog to digital conversion data can be transferred to the microcontroller via serial peripheral interface (SPI) bus.

The main computational element of the embedded system can be a 32-bit microcontroller. In the untethered operation state, the microcontroller performs the servo and activity controllers of the prosthesis and data logging at each sample time. In addition to untethered operation, the prosthesis can also be controlled via a tether by a laptop computer running MATLAB Simulink RealTime Workshop. In the tethered operation state, the microcontroller drives the servo amplifiers based on analog reference signals from the laptop computer. A memory card can be used for logging time-stamped data acquired from the sensors and recording internal controller information. The memory chip can be interfaced to the computer via wireless USB protocol. The microcontroller sends PWM reference signals to two four quadrant brushless DC motor drivers with regenerative capabilities in the second and forth quadrants of the velocity/torque curve.

As noted above with respect to FIG. 9, additional controls can be provided for operating the prosthesis when going from a sitting to a standing position or vice versa. This can be implemented via the use of a sitting mode controller implemented in the microcontroller. Operation of the sitting mode controller consists of four phases that are outlined in the general control state chart shown in FIG. 16. As shown in FIG. 16, two phases are primary sitting phases, weight bearing and non-weight bearing. The other two phases encompass the transition phases, pre-stand and pre-sit, for standing up and sitting down, respectively. Weight bearing and non-weight bearing are the primary sitting phases that switch the knee and ankle joints between high and low impedances, respectively. The transition phases, pre-stand and pre-sit, modulate the stiffness of the knee as a function of knee angle, as shown in FIG. 17, to assist the user in standing up and sitting down. FIG. 17 shows knee angle modulated knee stiffness during pre-stand (solid line) and pre-sit (dashed line) phases.

The modulation allows for smoother transitions near the seated position. The ankle joint can be slightly dorsiflexed with moderate stiffness during the standing up and sitting down phases. Switching between the four sitting phases occurs when sensor thresholds are exceeded, as depicted FIG. 16. The parameters of the impedance based controllers are tuned using a combination of feedback from the user and joint angle, torque and power data from the prosthesis.

In the various embodiments of the invention, actuation for a prosthesis can be provided by two motor-driven ball screw assemblies that drive the knee and ankle joints, respectively, through a slider-crank linkage. The prosthesis can be capable of 120° of flexion at the knee and 45° of planterflexion and 20° of dorsiflexion at the ankle. In one embodiment, each actuation unit consists of a DC motor (such as a Maxon EC30 Powermax) connected to a 12 mm diameter ball screw with 2 mm pitch, via helical shaft couplings. An exemplary ankle actuation unit additionally incorporates a 302 stainless steel spring (51 mm free length and 35 mm outer diameter), with 3 active coils and a stiffness of 385 N/cm in parallel with the ball screw.

As described above with respect to FIGS. 1A-4, the purpose of the spring can be to bias the motor's axial force output toward ankle plantarflexion, and to supplement power output during ankle push off. The stiffness of the spring can be maximized to allow for peak force output without limiting the range of motion at the ankle. The resulting axial actuation unit's force versus ankle angle plot can be shown in FIG. 18. FIG. 18 is a plot if axial force as a function of ankle angle illustrating spring force, actuator force and total force. FIG. 18 graphically demonstrates for fast walking the reduction in linear force output supplied by the motor at the ankle through the addition of the spring. Note that the compression spring does not engage until approximately five degrees of ankle plantarflexion. Each actuation unit can include a uniaxial load cell (such as Measurement Specialties ELPF-500L), positioned in series with the actuation unit for closed loop force control of the motor/ballscrew unit. Both the knee and ankle joints can incorporate bronze bearings and, for joint angle measurement, integrated precision potentiometers (such as an ALPS RDC503013). A strain based sagittal plane moment sensor, as previously described with respect to FIGS. 14A and 14B can be located between the knee joint and the socket connector, which measures the moment between the socket and prosthesis. The ankle joint connects to a foot, which incorporates strain gages to measure the ground reaction forces on the ball of the foot and on the heel. The central hollow structure houses a lithium-polymer battery and provides an attachment point for the embedded system hardware. To better fit with an anthropomorphic envelope, the ankle joint can be placed slightly anterior to the centerline of the central structure. This gives the prosthesis the illusion of flexion when the amputee can be standing vertically with the knee fully extended.

The length of the shank segment can be varied by changing the length of three components; the lower shank extension, the spring pull-down, and the coupler between the ball nut and ankle. Additional adjustability can be provided by the pyramid connector that can be integrated into the sagittal moment load cell for coupling the prosthesis to the socket (as is standard in commercial transfemoral prostheses).

Passive joint torque, τ_(p), can be defined as the part of the joint torque, τ, which can be represented using spring and dashpot constitutional relationships (passive impedance behavior). The system can only store or dissipate energy due to this component. The active part can be interpreted as the part which supplies energy to the system and the active joint torque can be defined as τ_(a)=τ−τ_(p). This active part can be represented as an algebraic function of the user input via the mechanical sensory interface (i.e. socket interface forces and torques).

Gait is considered a mainly periodic phenomena with the periods corresponding to the strides. Hence, the decomposition of a stride will give the required active and passive torque mappings for a specific activity mode. In general, the joint behavior exhibits varying active and passive behavior in each stride. Therefore, segmenting of the stride in several parts can be necessary. In this case, decomposition of the torque over the entire stride period requires the decomposition of the different segments and piecewise reconstruction of the entire segment period. In order to maintain passive behavior, however, the segments cannot be divided arbitrarily, but rather can only be segmented when the stored energy in the passive elastic element is zero. This requires that the phase space can only be segmented when the joint angle begins and ends at the same value. FIG. 19 shows the phase portrait of normal speed walking and the four different stride segments, S₁, S₂, S₃ and S₄. Thus, the entire decomposition process consists of first appropriate segmentation of the joint behavior, followed by the decomposition of each segment into its fundamental passive and active components.

The decomposition of each segment shown in FIG. 19 can be converted to an optimization problem. In each segment of the stride, 2n data points are selected by sampling the angular position in equal intervals between its minimum and maximum and selecting the corresponding positive and negative angular velocities. In this work, the number of angular position samples for each segment, n can be set to be 100. The constrained least squares optimization problem given in Equation 2 below can be constructed and solved.

$\begin{matrix} {{\min\limits_{x}{\frac{1}{2}{{{Cx} - d}}_{2}^{2}\mspace{14mu}{s.t.\mspace{14mu} 0}}} \leq x} & (2) \end{matrix}$ where C, x and d are defined in Equations 3, 4, and 5 below, respectively. The indexing of the joint angular position, angular velocity and moment samples are explained via the sketch in FIG. 20. FIG. 20 shows a selection and indexing of data samples from a first segment.

$\begin{matrix} {C_{4{nx}\; 3n} = \begin{bmatrix} C_{1} & C_{2} & C_{3} \end{bmatrix}^{T}} & (3) \\ {C_{1} = \left\lbrack {\begin{matrix} {{diag}\left( {\begin{bmatrix} \theta_{1} \\ \theta_{2} \\ \vdots \\ \theta_{n} \end{bmatrix}_{{nx}\; 1} - \alpha} \right)} \\ {{diag}\left( {\begin{bmatrix} \theta_{n} \\ \theta_{n - 1} \\ \vdots \\ \theta_{1} \end{bmatrix}_{{nx}\; 1} - \alpha} \right)} \end{matrix}{{diag}\left( \begin{bmatrix} {\overset{.}{\theta}}_{1} \\ {\overset{.}{\theta}}_{2} \\ \vdots \\ \vdots \\ \vdots \\ \vdots \\ {\overset{.}{\theta}}_{n} \end{bmatrix}_{2{nx}\; 1} \right)}} \right\rbrack_{2{nx}\; 3n}} & \; \\ {C_{2} = \left\lbrack {\begin{matrix} C_{21} \\ C_{22} \end{matrix}C_{23}} \right\rbrack_{{2n} - {1x\; 3n}}} & \; \\ {C_{21} = \begin{bmatrix} \theta_{1} & {- \theta_{2}} & 0 & \ldots & 0 \\ 0 & \ddots & \ddots & \ddots & \vdots \\ \vdots & \ddots & \theta_{n - 1} & \theta_{n} & 0 \\ 0 & \ldots & 0 & 0 & 0 \end{bmatrix}_{nxn}} & \; \\ {C_{22} = \begin{bmatrix} \theta_{n} & {- \theta_{n - 1}} & 0 & \ldots & 0 \\ 0 & \ddots & \ddots & \ddots & \vdots \\ \vdots & \ddots & \theta_{3} & {- \theta_{2}} & 0 \\ 0 & \ldots & 0 & \theta_{2} & {- \theta_{1}} \end{bmatrix}_{n - {1{xn}}}} & \; \\ {C_{23} = \begin{bmatrix} {\overset{.}{\theta}}_{1} & {- {\overset{.}{\theta}}_{2}} & 0 & \ldots & 0 \\ 0 & \ddots & \ddots & \ddots & \vdots \\ \vdots & \ddots & {\overset{.}{\theta}}_{{2n} - 2} & {- {\overset{.}{\theta}}_{{2n} - 1}} & 0 \\ 0 & \ldots & 0 & {\overset{.}{\theta}}_{{2n} - 1} & {- {\overset{.}{\theta}}_{2n}} \end{bmatrix}_{{2n} - {1x\; 2n}}} & \; \\ {C_{3} = \begin{bmatrix} \beta & \beta & \ldots & \ldots & \ldots & \beta & \beta \end{bmatrix}_{1x\; 3\; n}} & \; \\ {x_{3{nx}\; 1} = \begin{bmatrix} k_{1} \\ k_{2} \\ \vdots \\ k_{n - 1} \\ k_{n} \\ b_{1} \\ b_{2} \\ \vdots \\ b_{{2n} - 1} \\ b_{2n} \end{bmatrix}} & (4) \\ {d_{4{nx}\; 1} = \begin{bmatrix} \tau_{1} \\ \tau_{2} \\ \vdots \\ \tau_{{2n} - 1} \\ \tau_{2n} \\ {\tau_{1} - \tau_{2}} \\ {\tau_{2} - \tau_{3}} \\ \vdots \\ {\tau_{{2n} - 1} - \tau_{2n}} \\ 0 \end{bmatrix}} & (5) \end{matrix}$

The matrix C consists of three sub-matrices, C₁, C₂ and C₃. C₁ can be the main part responsible for the fitting of the spring and dashpot constants, k and b. C₂ bounds the rate of change of the passive joint torque and ensures smoothness in the resulting passive joint torque, and C₃ is basically a row of penalty constants, β, which penalizes large values of the spring and dashpot constants and thus limits the magnitudes of both. In this work, β is set to 0.1.

The origin of each virtual spring can be also added to the optimization problem formulation as a parameter in order to obtain a tighter passive torque fit. Therefore, the optimization problem given by (3) can be solved iteratively for a range of values of spring origin constant, α. The solution with the least error norm can be selected as the optimal solution.

The result of the above stated constrained optimization problem for segment 1 can be shown in FIGS. 21A, 21B, and 21C. FIGS. 21A, 21B, and 21C are the output of the decomposition for s₁ in FIG. 19 showing the spring and dashpot constants and the active and passive knee torques (Spring origin, α. is 23 degrees).

As can be seen from FIG. 21, the decomposed passive part can be very similar to the joint torque, and thus it can be stated that the behavior of the joint can be mainly passive. The result of the decomposition for the segment S_(i) can be stored in R_(i) of the form given in Equation 6. R _(i)=[θ {dot over (θ)} τ_(pas) F _(S1) F _(S2) τ_(act)]_(2n×6)  (6) where τ_(pas)=C₁x.

The procedure presented above decomposes the joint torques into active and passive parts. The joint torque references for the control of the prosthesis are generated by combining this active and passive torques. There are two major challenges to be solved. Firstly, the correct motion segment must be selected. Secondly, after the motion segment is selected at each sampling instant a new joint torque reference can be generated using the discrete mappings for the active and passive torque parts.

A switching system modeling approach incorporating both discrete and continuous states can be used for the reconstruction of the torque reference signal. The state chart shown in FIG. 22. will govern the discrete dynamics of the controller. Since the sequence of the segments can be ordered (i.e., the direction of the motion for a specific gait phase does not change), each segment can transition only to the next one, where the transition guard function can be written as a inequality in terms of θ and {dot over (θ)}. The transitions between segments take no time and the dynamics of the controller are governed by the {f_(p) _(i) (θ,{dot over (θ)}); f_(a) _(i) (F_(S))} pair at each sampling instant. The joint reference torque is τ_(ref)=τ_(a)+τ_(p) =f _(p) _(i) (θ,{dot over (θ)})+f _(a) _(i) (F _(S))  (7)

The decomposition algorithm presented above gives the result matrix, R, for each segment. The discrete data in R can be used to construct the joint torque reference for the continuous measurements of another trial in the same gait phase. At each sampling instant of the algorithm, the measurement vector m=[θ_(m),{dot over (θ)}_(m),F_(S1) _(_) _(m),F_(S2) _(_) _(m)]^(T) can be acquired. For the reconstruction of the passive knee torque part, the Euclidian error norm between the [θ_(m) {dot over (θ)}_(m)]^(T) and the angular position and velocities of all the samples in that segment [θ_(i) {dot over (θ)}_(i)]^(T) can be calculated as shown in Equation 8 and stored in the vector e. e _(i)=√{square root over ((θ_(m)−θ_(i))²+({dot over (θ)}_(m)−{dot over (θ)}_(i))²)}  (8) Then two elements of this vector with the least error norm are found and the passive knee torque reference can be found as a weighted linear combination of the passive knee torques corresponding to these points. The reconstruction of the active knee torque part is similar where only {θ,{dot over (θ)}, τ_(pas)} is exchanged with {F_(S1),F_(S2),τ_(act)}.

The supervisory controller (intent recognizer) switches among different underlying intramodal controllers depending on the activity mode the user imposes on the prosthesis. The intent recognizer consists of three parts: activity mode recognizer, cadence estimator and the slope estimator.

The activity mode recognizer detects the activity mode of the prosthesis (standing, walking, sitting, stair ascent or stair descent, etc. . . . ). This can be accomplished by comparing the features which are generated in real time to a feature database using some machine learning and/or pattern recognition methods. The present implementation of the gait mode recognizer, which recognizes standing and walking modes, is described below.

Firstly, a database which contains all the possible activity modes (standing and walking in this case) can be generated by making experimental trials. In the experimental trials, the user can be asked to walk or stand in different controller modes for 50 second long trials. The socket sagittal moment above the knee joint, foot heel load, foot ball load, knee angle, knee velocity, ankle angle and ankle velocity are recorded with 1 ms sampling period. It should be noted that other sensor signals such as accelerations and electromyography measurements from the residual limb can be added to the list of the signals used for intent recognition. For example, from the recorded experimental trials, 10000 random frames (5000 standing and 5000 walking) of 100 samples length are generated for all the seven recorded signals. The mean and the standard deviation of each frame are computed. The mean and standard deviation of signals are selected as the features since minimal computation can be required to obtain them. A database containing 10000 samples with 14 features (mean and standard deviation of the seven signals) belonging to two classes (standing and walking) can be generated. After the database is generated, the dimension of the database can be reduced from 14 to three using principal component analysis (PCA). Dimension reduction can be necessary because pattern recognition for high dimensional datasets can be computationally intensive for real-time applications. After dimension reduction step, the standing and walking data can be modeled with Gaussian mixture models. Gaussian mixture models represent a probability distribution as a sum of several normal Gaussian distributions. The order of the Gaussian mixture model for each mode can be determined according to the Minimum Description Length Criteria.

As described above, the database generation, dimension reduction and the Gaussian mixture modeling are explained. For real-time decision making, overlapping frames of 100 samples can be generated at each 10 ms interval. 14 features described above are extracted from these frames and the PCA dimension reduction can be applied to these features to get a reduced three dimensional feature vector. The reduced dimension features can be fed to the Gaussian mixture models for standing and walking and the probability of the sample vector being standing or walking can be computed. The mode with the greater probability is selected as the instantaneous activity mode. Since one decision might give wrong results in some cases due to noise, disturbance, etc. . . . , a voting scheme can be used to enhance the results. In the voting scheme, the controller activity mode is switched if and only if more than 90 percent of the instantaneous activity mode decisions among the last 40 decisions are a specific activity mode. Once a new activity mode is selected by the voting scheme, the underlying activity controller can be switched to the corresponding mode.

Such an activity mode recognizer is provided by way of illustration and not as a limitation. In the various embodiments of the invention, one or more parts of the algorithm might be modified. For example, in some embodiments, different features such as mean, max, kurtosis, median, AR coefficients, wavelet based features, frequency spectrum based features of the frame might be generated. Additionally, different dimension reduction techniques such as linear discriminant analysis, independent component analysis might be employed. Furthermore, different classification methods such as artificial neural networks, support vector machines, decision trees, hidden Markov models might be used.

Cadence estimation is accomplished by observing peak amplitudes in characteristic signal data and then measuring the time between successive peaks. Since walking is a cyclic activity each of the sensor signals will be periodic of cadence. The most relevant sensor signals will contain only one characteristic amplitude peak per stride such as foot heel load and the ball of foot load. In the real-time implementations, cadence estimation is accomplished by recording the foot load after heel strike when it exceeds 400 N until the load decreases below 350 N. Then, the time of occurrence of the peak load in this window is found and the previous peak time is subtracted from the new peak time. This corresponds to stride time and can be converted to cadence (steps/min) by multiplying with 120. Once the cadence is estimated, the intent recognizer selects the corresponding middle layer controller based on some predefined thresholds as in FIG. 23.

For example, in some embodiments, a 3D accelerometer capable of measuring ±3 g accelerations is embedded into the ankle joint coupler where the prosthetic foot is connected. An exemplary arrangement of such a system is shown by the schematic in FIG. 24. The accelerometer measurements are used to estimate the ground slope. In order to estimate the ground slope, the accelerometer data in tangential direction is used. Assuming the foot is flat on the ground, the ground slope angle, θ_(s), can be calculated as in equation (9) below.

$\begin{matrix} {\theta_{s} = {\sin^{- 1}\left( \frac{a_{t}}{g} \right)}} & (9) \end{matrix}$ In Eqn. 9, g is the gravitational constant. In order to find the ground slope estimate, {circumflex over (θ)}_(s), the accelerometer data should be collected while the foot is flat on the ground as determined by the heel and ball of the foot load sensors. While the foot is flat on the ground, equation (1) is computed for the frame of the collected data and the mean of this frame is outputted as the ground slope estimate, {circumflex over (θ)}_(s). Once the slope is estimated, the intent recognizer selects the corresponding middle layer controller based on some predefined thresholds. An exemplary state chart for such an intent recognizer is shown in FIG. 25.

Rather than a ballscrew and slider crank embodiment for the transmission of torque from a motor to the ankle and/or knee units, in some embodiments of the invention, the prosthesis can incorporate a friction and cable drive transmission embodiment. FIGS. 26A and 26B show front and back views of an exemplary embodiment of a friction drive transmission 2600 in accordance with an embodiment of the invention. As shown in FIGS. 26A and 26B, the shaft 2602 of an electric motor 2604 is preloaded against a first stage in a housing 2606, such as a larger diameter cylinder or friction drive gear 2608, which creates sufficient friction to transmit torque without slip. The shaft 2602 can use one or more friction rollers 2610 to transmit the torque. The first stage of the friction drive can also be supplemented with a second stage. The friction drive gear 2608 drives a smooth pinion 2612 directly, which is preloaded against a larger diameter cylinder or cable gear output 2614 in the housing 2606, which in turn transmits torque directly to the knee or ankle joint.

In addition to, or rather than a friction drive, the first or second stage of the transmission can alternatively be embodied by a cable drive transmission, in which a cable is wrapped around the circumference of a larger diameter cylinder, such as friction drive gear 2608, and also around the circumference of a smaller diameter cylinder, such as pinion 2612. In such embodiments, the cable is affixed to the friction drive gear 2608, and is pretensioned, using a tensioning screw 2616 or similar means, around both the drive gear 2608 and pinion 2612, such that friction between cable and pinion 2612 enables the transmission of torque from between the pinion 2612 and drive gear 2608. In one embodiment of a combined friction drive/cable drive transmission can be used, in which a first stage of the transmission (i.e., the friction drive gear 2608 connected directly to the electric motor 2604) is of the friction drive type, while the second stage of the transmission (i.e., the cable gear output 2614 connected directly to the knee or ankle joint) is of the cable drive type.

Rather than the ballscrew and slider crank or the friction drive and cable drive embodiments for the transmission of torque from a motor to the ankle and/or knee units, in some embodiments of the invention, the prosthesis can incorporate a chain drive or a belt drive transmission embodiment for implementing one or more stages of a transmission.

Advantages of a belt or chain drive approach over the ballscrew approaches described above include the ability to provide a fully enclosable/sealable (without need for a bellows-type cover) powered leg device. This facilitates component immersion in lubricating environment, and well as facilitating isolation from dirt, water, and other debris. As a result, this can extend the lifetime of transmission components. Another advantage of such a configuration is that it enables a greater range of motion of joint actuation, as opposed to a slider-crank mechanism (as used in a ballscrew configuration), which is generally limited. Further, the belt or chain drive approach also allows the device to maintain a constant transmission ratio throughout range of motion, which is not generally possible in the slider-crank mechanism typically used in a ballscrew configuration. Additionally, advantages of a belt or chain drive approach is that it maintains constant mechanism geometry throughout range of motion, belt and chain drive components are typically less expensive than ballscrew components, and belt and chain drive systems are typically characterized by lower audible noise than ballscrew configurations.

FIG. 27 shows an exemplary embodiment of a belt drive transmission 2700 in accordance with an embodiment of the invention. As shown in FIG. 27, a stage of the transmission 2700 can be embodied as a belt drive transmission, in which a belt 2702 is wrapped around the circumference of a larger diameter shaft, such as a first belt gear or pulley 2704, and also around the circumference of a smaller diameter shaft, such as second belt gear or pulley 2706. In such embodiments, the belt 2702 can be tensioned, using a tensioning device 2708. In one embodiment, the tensioning device 2708 can consist of a swing arm 2710, an additional pulley 2712 attached to the end of swing arm 2710, and tensioning screw 2714 for adjusting the swing arm 2710 to bias the additional pulley 2712 against the belt 2702, such that friction between the belt 2702 and belt gears 2704 and 2706 enables the transmission of torque from between second belt gear 2706 and first belt gear 2704. However, any other type of tensioning device can be used in the various embodiments to tension the belt 2702. For example, in some embodiments, the tensioning device 2708 can be a spring loaded device to automatically bias a pulley 2706 or other object against belt 2702 to cause the necessary tension.

It is worth noting that although transmission 2700 is illustrated in terms of a V-belt embodiment, the invention is not limited in this regard and can be used with any type of belts. For example, the belt 2702 can also be embodied as a flat belt, a round belt, a multi-groove belt, a ribbed belt, and a toothed or cog belt, to name a few. Further, the belt gears 2704 and 2706 can be configured in accordance with the type of belt being used.

In some embodiments, rather than utilizing a belt-based drive, a chain-based drive can be provided. The configuration in such embodiments can be substantially similar to that shown in FIG. 27. That is, a chain can be provided in place of belt 2702 and gears 2704 and 2706 can be embodied as sprockets compatible with the chain. In such embodiments, the tensioning device 2708 described above can still be utilized to maintain proper tension of the chain to enable the transmission of torque from between sprockets in the transmission.

In some embodiments, instead of utilizing a tensioning device as described above with FIG. 27, a pulley or sprocket can be configured with an eccentric mount. That is, configuring at least one of the drive gears in the transmission to allow an adjustment of its position. This is illustrated below with respect to FIGS. 28A-28D.

FIGS. 28A and 28B show side views of first and second positions, respectively, achievable for an exemplary embodiment of a chain drive transmission 2800 including an eccentric mount in accordance with an embodiment of the invention. Similar to the transmission described above with respect to FIG. 27, transmission 2800 includes a first shaft 2802 with first drive gears or sprockets 2804 and a second shaft 2806 with second drive gears or sprockets 2808 which can be coupled together via chains 2810 to transmit torques between sprockets 2804 and sprockets 2808. Although FIGS. 28A and 28B show that the transmission of torque between sprockets 2804 and sprockets 2808 is performed using two sets of sprockets (and thus using two chains), the embodiments are not limited in this regard. Rather, any number of chains can be used in the various embodiments.

As shown in FIGS. 28A and 28B, the first shaft 2802 is shown as including an additional sprocket 2812 for driving first shaft 2802. Such a configuration can be used when multiple drive stages are provided. However, the various embodiments are not limited in this regard.

In transmission 2800, the first shaft 2802 is configured to be eccentric. That is, the position of the first shaft 2802 is adjustable relative to the position of the second shaft 2806 so as to adjust the lateral separation between the shafts (i.e., to provide d_(A)≠d_(B)). Accordingly, this also provides a means to adjust the tension in a chain (or a belt) between the first shaft 2802 and the second shaft 2806. To provide the eccentric mount, the first shaft 2802 can be mounted in a leg device to an adjustable bearing mount 2814. The operation and configuration of an exemplary embodiment of the adjustable bearing mount 2814 is illustrated with respect to FIG. 29.

FIG. 29 illustrates schematically the components for the adjustable bearing mount 2812. As shown in FIG. 29, the adjustable bearing mount 2814 can include a top plate 2902 to which first shaft 2802 is attached, a bottom plate 2904, bearings 2906 between the top plate 2902 and the bottom plate 2904, and fasteners 2908. These components of the adjustable bearing mount 2814 can be disposed within an enclosure 2910.

In FIG. 29, the fasteners 2908 are shown as screws or bolts. However, the various embodiments are not limited to any particular bearing type or design of screws or bolts and other bearing types or designs can be used without limitation. Further, the various embodiments are not limited to screws or bolts and any other type of removable fastener can be used without limitation. Additionally, FIG. 29 shows bearings 2906 as a collection of ball bearings disposed between plates 2902 and 2904. However, the various embodiments are not limited to any particular bearing type or design and other bearing types or designs can be used without limitation.

In operation, the enclosure 2910 can be configured such that when fasteners 2908 are loosened or removed, the bearings allow the top plate 2902 can be repositioned relative to the bottom plate 2904 via bearings 2906. Thus, when fasteners 2908 are replaced and tightened, the plates 2902 and 2904 are biased against bearings 2906 to prevent further motion of the top plate 2902 relative to the bottom plate 2904.

Such a configuration allows adjustment of the position of first shaft 2802. For example, this can allow the first shaft 2802 to transition between a first position, as shown in FIG. 28A, in which a chain or belt 2810 with reduced tension is provided, due to a reduced distance (d_(A)) between first shaft 2802 and second shaft 2806, to a second position, as shown in FIG. 28B, in which a chain or belt 2814 with increased tension is provided, due to an increased distance (d_(B)) between first shaft 2802 and second shaft 2806. However, the various embodiments are not limited to solely first and second positions. Rather, in the various embodiments, the adjustable bear mount 2812 can be configured to allow a variety of positions for the first shaft 2806 relative to the second shaft 2806.

An exemplary configuration of a powered leg prosthesis 3000 in accordance with the discussion above is illustrated schematically in FIG. 30. As shown in FIG. 30, the powered leg prosthesis 3000 includes a shank 3002 with a powered knee joint 3004 and a powered ankle joint 3006. The powered knee joint 3004 includes a socket interface 3008 for attaching a socket 3010 or other device for attachment of the powered leg prosthesis 3000 to an amputee. The powered ankle joint 3006 can have a foot portion 3012 attached thereto.

The shank 3002 can consist of a single, discrete unit. However, in some embodiments, the shank can include an upper portion 3014 and a lower portion 3016. Such a configuration allows the insertion of at least one extension unit 3018 to allow the length of the shank 3002 to be customized for the amputee.

Within each of the upper portion 3014 and the lower portion 3016, a belt or chain drive system can be implemented, as described above with respect to FIGS. 27-29. For example, as shown in FIG. 30, the upper portion 3014 can include a first motor 3022, a first upper drive stage 3024, and a second upper drive stage 3026 for providing power at the powered knee joint 3004. Similarly, the lower portion 3016 can include a second motor 3028, a first upper drive stage 3030, and a second upper drive stage 3032 for providing power at the powered ankle joint 3006. Each stage can consist of the belt or chain drive stage. Additionally, each stage can be configured to include an eccentric mount, such as mounts 3034 and 3036, to adjust tension in the upper portion 3014 and lower portion 3016 respectively.

In addition to the components described above, the powered prosthetic leg 3000 can include other components not illustrated in FIG. 30 for purposes of clarity. For example, the powered prosthetic leg can include a control system or device, as previously described, and one or more sensors throughout the powered prosthetic leg, also as previously described. Thus control of the powered prosthetic leg 3000 can occur insubstantially the same manner as described above.

Coordinating Operation of Multiple Devices

As noted above, a first aspect of the various embodiments is the incorporation of a communications device in each of the autonomous prostheses in order to allow an autonomous prosthesis, such as a lower limb prosthesis, to be able to send data to or receive data from any other prosthesis associated with the user, such as a contralateral lower limb prosthesis. Such communications devices can be configured to form a communication network in which the communications devices communicate via wired or wireless communication links. Referring to FIG. 15 above, the wireless communications module can be used to provide such wireless communications between different prostheses. However, the various embodiments are not limited in this regard and additional components can be incorporated into each prostheses for supporting communications between the prostheses.

In operation, the prostheses can be configured to exchange various types of information. In particular, the prostheses can exchange configuration information in order to establish the overall configuration of the lower body, including, but not limited to, the configuration of both feet, shank, and thigh segments in both the frontal and sagittal planes. Thus, the configuration information can include ankle angle, knee angle, orientation of the respective segments in space, and a measurement of load bearing. In some embodiments, each of the prostheses can determine the state of the other based on such configuration information. However, the various embodiments are not limited in this regard and the information exchanged can also include the current control state of the respective prosthesis. Additionally, the information exchanged can also include perturbation information. That is, a perturbation detected in one prosthesis (e.g., a stumble event) can be communicated to the other prosthesis and used to recover from the perturbation.

A second aspect of the various embodiments is to configure the operation of each prosthesis to account for the state and configuration information in the data communicated thereto from another prosthesis. In particular, rather than relying solely on local data to determine how to operate, the prosthesis can be configured to utilize data from other prosthesis to infer information about the system it is operating in (i.e., how the body of a user is moving) and adjust the operation of the prosthesis based on this additional information. However, each prosthesis can continue to operate according to its own separate finite state model. That is, the prostheses are not in a master-slave relationship. Rather, the data from other prostheses are utilized solely to provide additional data to for the autonomous prosthesis to make an independent decision regarding its operation. Thus, the prostheses only provide each other with an augmented set of information regarding the environment and the system they are operating in (i.e., the amputee). Thus, the data from other prostheses can be used to constrain operation of prosthesis. As a result, the control systems of the various prostheses effectively coordinate, yet continue to operate independently.

In one exemplary configuration, the controller for each of the prostheses can be configured as follows. First, the non-weight bearing states of either prosthesis are constrained to occur only during weight-bearing states of the contralateral prosthesis. Thus, this ensures that at least one of the prostheses is supporting the user and reducing the likelihood of injury resulting from a fall. Second, in the case of devices capable of power delivery, the introduction of energy into the prosthesis by the controller is constrained to depend upon the communicated state of the contralateral prosthesis.

Therefore, although each of the prostheses independently adds energy as determined by their corresponding control systems, this energy delivery can be dictated by the behavior or state of the contralateral prosthesis. Therefore the energy added by a prosthesis is made dependent upon appropriate conditions being reported in the other prosthesis. An example of such a constraint would be limiting a transition to ankle push off in level ground walking to depend upon the contralateral prosthesis being configured for a heel strike. Such a configuration is provided to ensure a stable transition as the user shifts his or her weight from the ipsilateral prosthesis to the contralateral prosthesis.

Additionally, the controller can also be configured to respond to certain circumstances. For example, when the controller for a first leg receives information that a specific event has occurred or has been detected at the second leg, the controller can take appropriate action at the second leg to avoid injury of the amputee based on the current state of the second leg. For example, the controller can operate in accordance with a stumble recovery methodology, such as that described in pending U.S. patent application Ser. No. 13/508,175, entitled “Systems and Control Methodologies for Improving Stability in Powered Lower Limb Devices” and filed May 4, 2012, the contents of which are herein incorporated by reference in their entirety.

An advantage of the various embodiments is that the communication allows the amputee to perform activities of daily living, including level ground and sloped walking, stair ascent and descent, and running, and standing on uneven terrain with a greater degree of confidence. In particular, since essentially two sets of sensors are provided (one for each prosthesis), the data from each set can be combined to infer information regarding the terrain. For instance, in order to determine the ground slope or step height in a double stance configuration of the prostheses, the configuration of the joints of each prosthesis can be combined. This combined information can provide redundancy in a ground slope measurement based on inertial sensors and the angle of the ankle. This redundancy can be used to minimize error in the ground slope estimate and better select a control strategy for the terrain conditions. In addition, the configuration of both lower limbs in a double stance situation will indicate a discrepancy in ground height between the two prosthetic feet. In the condition where the leading limb has a more flexed configuration than the trailing limb, the leading limb is encountering an elevated step and a stair ascent strategy can be selected for both prostheses. In the condition where the leading limb has a less flexed configuration than the trailing limb, the leading limb is encountering a lower step and a stair descent strategy can be selected.

Another aspect of the various embodiments is that in the event of a breakdown in communication between the prostheses, each of the prostheses can be configured to revert to an alternate control methodology that maintains all feasible functionality achievable without contralateral information. Since all lower limb activities described in this document have previously been implemented in a unilateral control system for unilateral amputees, these activities would retain the same functionality as is present in the unilateral system. The loss of communication would only remove the enhancements made with respect to estimation of the terrain configuration as mentioned above, along with the safety enhancements such as preventing both prostheses from entering non-weight bearing conditions simultaneously.

In some embodiments, the communication between the prostheses can also be exploited to assist in the performance of certain types of tasks. For example, in some configurations, the communication between the prostheses can be utilized to achieve coordinated sit-to-stand and stand-to-sit transitions. There is some evidence that bilateral transfemoral amputees prefer asymmetric techniques to sit-to-stand and stand-to-sit transitions. One such technique for stand-to-sit is to bend the knee of one prosthesis first, while the contralateral knee remains extended. This is achieved by rotating the torso as it descends. In order to accommodate this motion, the leading (bending) prosthesis must first ensure that the contralateral prosthesis is extended. Furthermore, the trailing prosthesis must know that the contralateral prosthesis has begun to descend so that it, in turn, can change its behavior to knee flexion at the right instance. Such an asymmetric coordination would not be possible without communication between the prostheses.

During some types of activities, it may be useful for one or both prostheses of a bilateral amputee to enter a low-power state. For example, while sitting for extended periods of time it would be inefficient to actively control the joints of the prostheses, as they are largely unused in this configuration. If the user would like to stand, however, he or she may prefer an asymmetric sit-to-stand technique (in a manner similar to the stand-to-sit technique described above). In this case, the leading prosthesis can wake the contralateral prosthesis so that they are both ready to support the user for the sit-to-stand transition. In such cases, either leg can wake the other leg from a low-power state based on input from sensors.

EXEMPLARY RESULTS

The following examples and results are presented solely for illustrating the various embodiments and are not intended to limit the various embodiments in any way.

A. Mechanical Design

A pair of new prosthesis prototypes was constructed, similar to those described above with respect to FIGS. 1A-30. The frames of the prostheses are constructed from CNC-milled 7075 aluminum alloy. Each joint is actuated by a 30 mm Maxon EC 4-pole brushless DC motor through a back-drivable transmission with a ratio of 176:1 for the knee joints and 197:1 for the ankle joints. The prosthetic feet are constructed from carbon-fiber and are custom manufactured by Freedom Innovations, LLC. The feet are contained within prosthetic foot shells which allow for the use of standard shoes. The prostheses connect to the user's socket with a standard pyramid adapter. The upper portion of the shank houses an 8 cell, 30 V lithium-polymer battery with a capacity of 3300 mA·h.

Each prosthesis was configured to measure a group of mechanical signals for use in the control system: the angular positions of each joint, the angular velocities of each joint, and the axial load in the shank. A 6-axis inertial measurement unit is also included on the printed circuit board. The build height of the prosthesis is configurable via a bridge connector that attaches the upper shank segment to the lower shank segment. The bridge connector can be manufactured in different heights to accommodate a range of shank lengths. The minimum build height of the prosthesis (i.e. with the bridge connector completely removed) is 40 cm (15.75 in) from the knee center to the ankle center. This build height places the knee 45 cm (17.72 in) from the bottom of the sole plate, with actual knee height depending upon the foot shell and shoe selected by the user.

B. Embedded System

Each powered prosthesis was controlled with an embedded system that contains two microcontrollers. A 16-bit digital signal processor from Microchip (dsPIC33FJ64GS608) drives two custom 4-quadrant brushless servo-amplifiers to control current to each motor. A 32-bit general purpose microcontroller from Microchip (PIC32MX575F512L) performs the high level control activities (such as generating a current reference for the dsPIC), along with system overhead (data logging, power management, etc.) at a sampling rate of 500 Hz.

Communication between the prostheses is accomplished via a Controller Area Network (CAN) hardware peripheral module included in the PIC32 microcontroller. This module implements the data link layer and most of the physical layer, although the CAN specification itself does not dictate the physical medium for communication. The CAN protocol controller interfaces with a 3.3 V, 1 Mbps CAN Transceiver (MAX3051) from Maxim. The transceiver drives a 250 cm shielded twisted pair data tether that connects to the contralateral prosthesis. In normal operation the data tether is routed along the user's thighs and waist, though future work will involve moving to a wireless interface.

The CAN module is configured to run at its maximum speed of 1 Mbps. Using the standard CAN data frame of 111 bits, the maximum possible frame speed is approximately 9446.63 frames per second. For a sampling period of 2 ms (for the sample rate of 500 Hz for the main control loop), the CAN bus could handle no more than 18 complete frames during each sample. For a symmetric control scheme each prosthesis could then ideally send 9 frames of data per sample time, assuming no collisions occurred on the bus. In order to safeguard against an overrun due to collisions on the bus, each prosthesis only attempts to send 4 frames per sample time. Each frame contains two 32-bit integers of data, so eight 32-bit variables can be transferred in each direction at the sampling rate of 500 Hz. The control variables that are shared between the legs can be seen in Table III.

TABLE I VARIABLES SHARED OVER CAN BUS Frame Variable Name Description 1 run_time an internal counter that increments every sampling instant control_modes a 32-bit structure that contains the current states of the state machines 2 knee_position the angular position of the prosthetic knee joint knee_velocity the angular velocity of the prosthetic knee joint 3 ankle_position the angular position of the prosthetic ankle joint ankle_velocity the angular velocity of the prosthetic ankle joint 4 shank_load the axial load of the shank unimplemented reserved for future implementation

C. Impedance Control

The servo-amplifier in the embedded system measures and controls the current to the brushless DC motors driving each joint. A simple feed-forward model was implemented to transform desired torques into current references for the servo-amplifier. The transmission ratio is fixed by the design of the device, the torque constant is provided by the motor manufacturer, and the Coulomb friction estimate is empirically derived for the prosthesis. The torque commands are generated by the high level controller according to the following impedance control law specified in Eq. 10. τ_(k) =k _(k)(θ_(k)−θ_(eq) _(k) )+b _(k){dot over (θ)}_(k) τ_(a) =k _(a)(θ_(a)−θ_(eq) _(a) )+b _(a){dot over (θ)}_(a)  (10) In Eq. 10, the torque reference (τ_(k) for the knee, τ_(a) for the ankle) is a linear function of the position and velocity of the joint (θ_(k) and {dot over (θ)}_(k), respectively, for the knee, and θ_(a) and {dot over (θ)}_(a) for the ankle). The three parameters k, b, and θ_(eq), and specify the impedance (stiffness, damping coefficient, and equilibrium position) of the joint and are software selectable for each state in the state machine.

D. Finite State Machine for Level Walking

The level walking gait controller is implemented as a finite state machine where each state corresponds to a unique impedance, as specified by the impedance parameters k, b, and θ_(eq) for both the knee and the ankle. The form of the controller follows the unilateral prosthesis controller, although several transitions have been enhanced using the CAN tether. Most notably, the state of the contralateral prosthesis is now considered during transitions into and out of the push off state. This is illustrated with respect to FIGS. 31 and 32. FIG. 31 shows exemplary state chart for bilateral, autonomous powered leg prostheses. FIG. 32 illustrates the communication and coordination between two autonomous powered leg prostheses, each using the controller of FIG. 31, based on establishment of a controller area network (CAN) in accordance with the various embodiments.

The state transitions for the controller in FIG. 31 are listed in Table II.

TABLE II STATE TRANSITIONS OF THE WALKING CONTROLLER Transition Description Thresholds T01 Foot has fully landed in stance. ↑ Ankle torque T12 Ankle begins actively pushing off. ↑ Ankle Angle T22 Knee angle impedance lowered ↑ Knee Angle T23 Foot leaves the ground. ↓ Shank Load T34 Maximum knee flexed is reached. ↓ Knee Velocity T45 Knee straightens before heel strike. ↓ Knee Angle T50 40 30 Ground contact is detected. ↑ Shank Load

As shown in FIGS. 31 and 32, both prostheses begin in the middle stance state (1) before gait is initiated. In middle stance, the knee and ankle are high impedance and the prosthesis behaves as a stiff spring, supporting the user. The user can begin the gait cycle by stepping forward with one prosthesis. At this time, the virtual spring at the trailing ankle allows dorsiflexion as the user shifts more weight to the stance limb. Once an angle threshold is met in the ankle (T12), the prosthesis transitions into push off (2), provided that the contralateral prosthesis is not already in the push off state based on state information from the contralateral prosthesis. This safety condition prevents both prostheses from simultaneously pushing off, as otherwise would be the case if both ankles underwent dorsiflexion at once (i.e. in the case of leaning forward while standing).

During push off (2) the knee behavior switches from high impedance (2A) to low impedance (2B) once the knee angle passes its equilibrium position (T22). This transition allows the powered push off from the ankle to impart energy to the user's center of mass initially and then subsequently to impart energy to the knee for swing. In the low impedance portion of push off it is imperative that the user is in double support, as the user's weight is transitioning to the contralateral limb. Therefore, the knee impedance transition during push off is contingent upon the contralateral prosthesis being in a weight-bearing state (i.e. either heel strike or middle stance). Again, this safety condition prevents both prostheses from simultaneously pushing off, as otherwise would be the case if both ankles underwent dorsiflexion at once (i.e. in the case of leaning forward while standing).

Once ground contact is no longer detected (T23) by the prosthesis (as measured by the load cell), it transitions into swing, which consists of a flexion state (3), an extension state (4), and a final state, called pre-landing (5), where the knee adopts a moderate impedance in preparation for heel strike (0). Heel strike (0), which occurs in response to detecting a shank load (T50) consists of a stiff knee and a moderately damped ankle for shock absorption, and then quickly transitions into middle stance for the next stride. Each of the other swing states can also transition to heel strike if a load (as measured by the load cell) is detected during swing flexion (T30) or swing extension (T40). This safety condition causes the prosthesis to transition to a state that provides support to the user in the case of a stumble or otherwise premature contact with the ground.

E. Able Body Adapters

In order to assess the feasibility of using the powered prostheses for gait in bilateral amputee subjects, a pair of able body adapters was constructed such that a healthy subject could don the prostheses. The adapters were constructed from aluminum stock, plastic and commercially available sports equipment. The adapters were secured with hook-and-loop straps and immobilized the healthy knee joint at approximately 100 degrees of flexion. The prosthetic knee joint was located below the adapters and placed posterior to the healthy joint in order to comfortably assume the subject's body weight.

The subject practiced using the adapters in conjunction with the powered prostheses and also a pair of passive prostheses. For the passive prosthesis the subject used the Ossur Total Knee 2000 in conjunction with the Pacifica LP carbon fiber ankle-foot complex from Freedom Innovations.

Impedance parameters for the powered prostheses were tuned experimentally based upon a data set used in previous studies. The impedance parameters were tuned identically for each leg, and are listed in Table III.

TABLE III IMPEDANCE PARAMETERS FOR THE POWERED PROSTHESIS Knee k Knee b Knee Ankle k Ankle b Ankle Finite Nm/ Nm · s/ θ_(eq) Nm/ Nm · s/ θ_(eq) State deg deg deg deg deg deg Heel 3.5 0.15 8.0 5.0 0.3 0 Strike Middle 5.0 0.35 10.0 6.0 0.15 −3.0 Stance Push Off: 4.0 0.025 15.0 5.0 0.15 −15.0 Stiff Knee Push Off: 0.3 0.025 15.0 5.0 0.15 −15.0 Soft Knee Swing 0.3 0 40.0 1.5 0.15 0 Flexion Swing 0.3 0.015 30.0 2.0 0.25 0 Extension Pre- 0.2 0.15 15.0 2.5 0.25 0 Landing

F. Treadmill Setup

Parallel bars were constructed around a treadmill in a motion capture laboratory in order to conduct the gait experiments. The subject was allowed to use the parallel bars for stability through the tests. The subject walked at treadmill speeds of 2.0 mph, 2.3 mph, and 2.5 mph on both pairs of prostheses. Motion capture data was collected at 120 Hz through 12 infrared cameras (model S250e from Optitrack) and processed in NaturalPoint's ARENA software environment. Single axis joint angles (sagittal plane) for the hips, knees, and ankles were extracted from the 3D skeleton data using MATLAB.

G. Results

The subject was able to walk at all three speeds using both pairs of prostheses. Sustaining a speed of 2.5 mph with the passive prostheses, however, was difficult, as the prosthetic knees did not allow a fast enough extension in swing to match the subject's cadence. For each condition (treadmill speed and prosthesis type) 15 consecutive strides were selected for analysis from the middle of the trial. However, for the case of passive devices at 2.5 mph, only 8 consecutive strides were achieved in steady state conditions.

FIG. 33 displays the averaged kinematics of the hips, knees, and ankles for a group of healthy subjects from the literature, the 15 strides at each speed for the powered trials, and the 15 strides (8 for 2.5 mph) at each speed for the passive trials. The healthy subject data from the literature did not explicitly state speed or cadence, but referred to natural, slow and fast cadences (estimated to be between 100-120 steps/min). It should be noted that all three of these cadences are faster than the fastest cadence achieved by the subject.

H. Discussion

The hip kinematics for both the passive and powered prostheses differ from the healthy subject data from the literature. However, since kinetics were not measured in this preliminary work, we cannot yet definitively account for these differences. In both cases the hips are biased 10 degrees at the beginning of stance and 20 degrees just after peak knee flexion in swing. This offset, however, is not (directly) due to the biasing of the thighs with the able body adapters, as the effective thigh angle was calibrated in order to remove this offset statically.

The knee kinematics of the powered prostheses show a marked improvement over the passive prostheses. The powered data shows a slightly diminished but comparable amount of knee flexion in stance (from 0 to 40% of stride), while the passive knee, which is a four-bar mechanism, must be locked to provide support in stance. Consequently, stance knee flexion is absent in the passive knee kinematics. The powered swing phase, as determined by the knee flexion in the second half of the stride, more closely matches the literature than that of the passive knee. A shortened swing indicates that a greater percentage of the stride is spent in double support, which may be due to a lack of stability on the passive devices.

The ankle kinematics of the powered prostheses closely matched that of the healthy subjects from the literature. A period of energy absorption is present following heel strike as foot-flat is reached. Evidence of a powered push off can be seen in the extreme plantarflexion of the ankle at approximately 60% of stride. The only significant deviation from the healthy data is the rate of ankle dorsiflexion in stance. The healthy ankle appears to mimic a nonlinear spring in stance, while the powered prosthesis is using a linear impedance model. The result is a lag in dorsiflexion at the beginning of stance (the emulated spring is too stiff), and too rapid of a dorsiflexion at terminal stance (the emulated spring is too soft). This phenomenon is echoed in the stance knee flexion, which is slightly less than that seen in the literature. The passive ankle shows decent behavior as well, though this particular ankle appears to be too soft for the user, as it shows a large amount of compliance at heel strike and a comparably large amount of dorsiflexion in late stance. These deviations could be reduced by using a stiffer ankle, though the lack of energy transfer at push off, as evidenced by the lack of a large plantarflexion around 60% of stride, is an inherent limitation of the passive device.

While various embodiments of the present invention have been described above, it should be understood that they have been presented by way of example only, and not limitation. Numerous changes to the disclosed embodiments can be made in accordance with the disclosure herein without departing from the spirit or scope of the invention. Thus, the breadth and scope of the present invention should not be limited by any of the above described embodiments. Rather, the scope of the invention should be defined in accordance with the following claims and their equivalents.

Although the invention has been illustrated and described with respect to one or more implementations, equivalent alterations and modifications will occur to others skilled in the art upon the reading and understanding of this specification and the annexed drawings. In addition, while a particular feature of the invention may have been disclosed with respect to only one of several implementations, such feature may be combined with one or more other features of the other implementations as may be desired and advantageous for any given or particular application.

The terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting of the invention. As used herein, the singular forms “a”, “an” and “the” are intended to include the plural forms as well, unless the context clearly indicates otherwise. Furthermore, to the extent that the terms “including”, “includes”, “having”, “has”, “with”, or variants thereof are used in either the detailed description and/or the claims, such terms are intended to be inclusive in a manner similar to the term “comprising.”

Unless otherwise defined, all terms (including technical and scientific terms) used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. It will be further understood that terms, such as those defined in commonly used dictionaries, should be interpreted as having a meaning that is consistent with their meaning in the context of the relevant art and will not be interpreted in an idealized or overly formal sense unless expressly so defined herein. 

What is claimed is:
 1. A method for operating an autonomous powered leg prosthesis for a lower limb of a body having at least one powered joint, at least one sensor for collecting real-time sensor information for the autonomous powered leg prosthesis, and a communications device, the method comprising: generating, via a first control system of the autonomous leg prosthesis, first control data for the communications device to establish a communications link with a second leg prosthesis associated with another lower limb of the body contralateral to the lower limb and having a second control system operating independently of the first control system; receiving, at the first control system over the communications link, remote data for the second leg prosthesis, the remote data comprising state information for a second finite state model used by the second control system for operating the second leg prosthesis; generating, via the first control system, second control data for transitioning the autonomous leg prosthesis from a current state in a first finite state model for operating the autonomous powered leg prosthesis to a different state in the first finite state model, wherein the different state is selected based on the real-time sensor information and the remote data, wherein at least one state in the first finite state model is configured for causing the at least one powered joint to deliver net power, wherein the second control data further comprises instructions to transition the autonomous leg prosthesis to the at least one state only if the state data in the remote data indicates that the second leg prosthesis is in a weight bearing state.
 2. The method of claim 1, wherein the remote data further includes real-time sensor information for the second leg prosthesis.
 3. The method of claim 1, wherein the generating of the second control data further comprises: determining a local terrain configuration based on the real-time sensor information and the remote data, and configuring the first control data based on the local terrain.
 4. The method of claim 1, wherein the generating of the second control data further comprises: detecting a perturbation based on the remote data, and configuring the first control data to compensate for the perturbation.
 5. The method of claim 1, further comprising detecting a failure in the communications link, and wherein the generating of the second control data further comprises configuring the second control data to cause the autonomous powered leg prosthesis to transition to one of one or more pre-defined safe states in response to the failure.
 6. The method of claim 1, wherein the generating of the second control data further comprises: detecting one of a sit-to-stand transition and a stand-to-sit transition based on the real-time sensor information and the remote data to yield a detected transition, and configuring the second control data based on the detected transition.
 7. A non-transitory computer-readable medium having stored thereon a computer program executable by a first control system of an autonomous powered leg prosthesis for a lower limb of a body having at least one powered joint, at least one sensor for collecting real-time sensor information for the autonomous powered leg prosthesis, and a communications device, the computer program comprising a plurality of code sections for causing the first control system to perform the steps of: generating first control data for the communications device to establish a communications link with a second leg prosthesis associated with another lower limb of the body contralateral to the lower limb and having a second control system operating independently of the first control system; receiving, over the communications link, remote data for the second leg prosthesis, the remote data comprising state information for a second finite state model used by the second control system for operating the second leg prosthesis; generating second control data for transitioning the autonomous powered leg prosthesis from a current state in a first finite state model for operating the autonomous powered leg prosthesis to a different state in the first finite state model, wherein the different state is selected based on the real-time sensor information and the remote data, wherein at least one state in the first finite state model is configured for causing the at least one powered joint to deliver net power, wherein the second control data further comprises instructions to transition the autonomous leg prosthesis to the at least one state only if the state data in the remote data indicates that the second leg prosthesis is in a weight bearing state.
 8. The non-transitory computer-readable medium of claim 7, wherein the remote data further includes real-time sensor information for the second leg prosthesis.
 9. The non-transitory computer-readable medium of claim 7, wherein the generating of the second control data further comprises: determining a local terrain configuration based on the real-time sensor information and the remote data, and configuring the first control data based on the local terrain.
 10. The non-transitory computer-readable medium of claim 7, wherein the generating of the second control data further comprises: detecting a perturbation based on the remote data, and configuring the first control data to compensate for the perturbation.
 11. The non-transitory computer-readable medium of claim 7, further comprising detecting a failure in the communications link, and wherein the generating of the second control data further comprises configuring the second control data to cause the autonomous powered leg prosthesis to transition to one of one or more pre-defined safe states in response to the failure.
 12. The non-transitory computer-readable medium of claim 7, wherein the generating of the second control data further comprises: detecting one of a sit-to-stand transition and a stand-to-sit transition based on the real-time sensor information and the remote data to yield a detected transition, and configuring the second control data based on the detected transition. 